Stents are used to treat atherosclerotic stenosis or other type of blockages in body lumen like blood vessels or to expand the lumen that has narrowed due to disease. “Stenosis” is narrowing of the diameter of a bodily passage or orifice due to formation of plaque or lesion. The function of the stent is to expand the lumen diameter by pressing the plaque to the vessel wall and to maintain patency of the lumen of the blood vessel thereafter at the location of its implantation. The stent may be coated with therapeutic agent/s and/or biocompatible material/s for beneficial effects like minimizing the possibility of restenosis, reduction in inflammation etc.
The first step in treatment of stenosis involves locating the region that may require treatment such as a suspected lesion in a vessel by angiography of the diseased vessel followed by implanting a suitable stent. The stent may be balloon expandable type or self-expanding type. The stents are mounted on the delivery catheter which helps in delivering the stent to the target site of the disease.
The balloon expandable stent is mounted on a balloon catheter by crimping process such that it holds tightly over the balloon and attains a considerably lower diameter (profile). The catheter is percutaneously inserted into the body lumen and is directed to the site of the disease (blockage or narrowed lumen). At the site of the disease, the balloon is inflated by application of hydraulic pressure to expand the stent radially to desired diameter. Radial expansion of the stent presses the plaque to the wall of the vessel by which the restriction to the flow of blood in the vessel is removed. The balloon is then deflated by removing the hydraulic pressure and withdrawn from the body of the patient.
On expansion, the stent material attains plastic deformation and hence the stent does not recoil back to its original shape and remains in expanded state keeping the lumen patent. Self-expanding stents are typically made of metal with shape memory and they expand without help of any other device such as balloon. The stent is mounted on the delivery catheter and expansion of the stent is restricted by a sheath. The catheter is percutaneously inserted into the body lumen and guided to the target site where the lesion or plaque is located. The sheath is then retracted to allow the stent to expand. Like balloon expandable stent, this stent also makes the lumen patent by pressing the plaque. The structure of stents is cylindrical with scaffold made up of a pattern or network of interconnecting structural elements i.e. “struts”. The scaffolding of the stent may be formed from wires, tubes, or sheets of material rolled into a cylindrical shape. In addition, the surface of the stent may be coated with formulation of therapeutic agent/s and/or biocompatible materials with suitable carriers and additives.
It is important that the stent must be able to withstand structural loads viz. radial compressive forces imposed by the wall of the body lumen on the stent. Radially directed force from the wall of the lumen may tend to cause the stent to recoil inward. The radial strength of the stent must be adequate to resist radial compressive forces. These forces are cyclic in nature due to pulsating blood flow. Hence the stent should have adequate fatigue strength to withstand cyclic loading imposed on it by the lumen. In addition, the stent must possess sufficient flexibility to allow for crimping, maneuvering through the vascular pathway and expansion process. The scaffold structure should also be dense enough to prevent prolapse of the plaque but open enough to allow easy side branch access for another catheter with or without stent. The stent should exhibit required radio opacity for ease of implantation.
Stents have been used effectively for quite a long time and the safety and efficacy of stenting procedure are well established. Implantation of the stent causes some injury to the vessel. The healing process starts and finally the endothelial cells are formed at the implantation site. Once the healing process is completed the endothelial cells provide sufficient support to the wall of the lumen and the stent is no longer required. Thus, the presence of the stent in the lumen is required only for a limited period of time till the healing process is completed.
Coronary stents are generally made from biocompatible materials such as metals which are bio-stable. Metal has high mechanical strength that provides adequate radial and fatigue strength to the stent that prevent early and later recoil. However, the metallic stent remains at the implant site indefinitely. Leaving the stent at the implanted site permanently causes compliance difference in the stented segment and the healthy vessel segment. In addition, there is a possibility of permanent interaction between the stent and the surrounding tissue resulting in a risk of endothelial dysfunction causing delayed healing and late thrombosis.
Drug-eluting stents are a breakthrough in the development of stents with their ability to significantly reduce restenosis rates and the need for repeat revascularization. However, they are still associated with sub-acute and late thrombosis that necessitates prolonged antiplatelet therapy for at least 12 months.
Metallic stents have been used effectively for quite a long time and their safety and efficacy are well established. The main issues of a stent are restenosis and in-stent thrombosis. One of the important causes of these adverse effects is injury of the artery caused by implantation of the stent. The injury leads to restenosis and delayed endothelialization. These adverse effects can be reduced if injury to the artery is reduced. It is well established that the thickness of the struts of a stent plays an important role in injuring of the artery. Thinner struts cause less injury compared to thicker struts. Thus, the injury of the artery can be reduced by making the struts as thin as practically possible.
While deciding the thickness of the struts, care should be taken so that important mechanical properties of the stent like radial strength and fatigue resistance are adequate to withstand forces imposed by the body lumen like artery.
The injury to the artery wall can thus be minimized by reducing the thickness of the struts of the stent scaffold structure. It is well established that the stent with less strut thickness causes less injury compared to the stent with thicker struts. This subject is discussed in detail by Kastrati A, Schomig A, Dirschinger J, et al. in their paper “Strut Thickness Effect on Restenosis Outcome (ISAR STEREO Trial)” published in Circulation 2001; 103:2816-2821. The incidence of angiographic restenosis was 15.0% in the group of patients treated with stents of thin struts against restenosis of 25.8% in the group treated with stents with thicker struts. Clinical restenosis was also significantly reduced, with a reintervention rate of 8.6% among thin-strut patients and 13.8% among thick-strut patients.
These findings were reconfirmed by Kastrati A, et al in their paper “Strut Thickness Effect on Restenosis Outcome (ISAR STEREO-2 Trial)” published in J. Am. Coll. Cardiol, 2003; 41:1283-8. The incidence of angiographic restenosis was 17.9% in the group of patients treated with stents of thin struts against restenosis of 31.4% in the group treated with stents with thicker struts. Target Vessel Revascularization (TVR) due to restenosis was required in 12.3% of the patients in thin strut group against 21.9% required in patients of the thick strut group.
In conclusion from above it was established that the use of thinner strut device is associated with a significant reduction in angiographic and clinical restenosis after coronary stenting. The stents can be made from polymeric materials which are bio-absorbable/biodegradable.
A biodegradable stent can be configured to degrade and disappear from the implant site when it is no longer needed leaving behind only the healed natural vessel. This will allow restoration of vasoreactivity with the potential of vessel remodeling. These stents are believed to improve the healing process whereby the chances of late stent thrombosis are reduced considerably. Prolonged antiplatelet therapy then may not be necessary. The biodegradable stents may be made from biocompatible polymers such as Poly-L-lactic acid (PLLA), polyglycolic acid (PGA), poly (D, L-lactide/glycolide) copolymer (PDLA), and polycaprolactone (PCL). Poly-L-lactic acid (PLLA) is usually recommended polymer among others.
The only disadvantage of polymeric materials is their lower mechanical strength compared to the metals. Strength to weight ratio of a polymeric material is smaller than that of a metal. This makes it necessary to increase thickness of the polymeric stents compared to metallic stents to get adequate radial and fatigue strengths. The increase in thickness results into higher profile and higher degree of injury to the blood vessel. Higher thickness reduces the flexibility of the stent resulting into poor trackability through tortuous arteries. Polymeric materials have poor radio opacity. Polymeric materials are also brittle under conditions within human body.
It is hence necessary to select right polymer and modify its mechanical properties to make it suitable for stent application. Making the stent with low strut thickness poses additional challenge. Selection of a polymeric material, design of stent scaffold structure and process for making a stent require careful attention to several aspects. The stent should have adequate mechanical strength to prevent recoil. The rate of degradation of the polymer should be such that the mechanical strength of the stent is retained to provide support to the vessel and prevent prolapse of the plaque into the vessel till the healing process is complete. The stent should eventually disappear by degradation. The stent should have enough flexibility for ease of crimping on the balloon of the catheter and for good trackability through the tortuous passages through arteries. The polymeric material and its degradation products should be biocompatible. The rate of degradation will influence the release profile of the therapeutic agents coated on the stent. Polymers such as Poly-L-lactic acid (PLLA), polyglycolic acid (PGA), poly (D, L-lactide/glycolide) copolymer (PDLA), and polycaprolactone (PCL) and their degradation products are known to be non-toxic and biocompatible.
There is a continuing need for manufacturing and fabricating methods for polymeric stents with such scaffold design that offer adequate radial strength, fracture toughness, low recoil and sufficient shape stability with low strut thickness. A stent with low strut thickness will result in low injury to the arterial wall. In addition, thin stent will give lower profile in crimped condition compared to a stent with higher strut thickness. Stents with thinner struts impart more flexibility to the stent. There is ample literature available on biodegradable stents and the process for manufacturing the same.
U.S. Pat. No. 7,971,333 describes a method of forming a stent from polymeric materials by modifying the mechanical properties of polymer tube to get desirable mechanical properties. The polymer can be modified to increase the strength, modulus and/or toughness of the polymer tube to make them comparable to metal. Mechanical properties of a polymer can be modified by applying stress to the polymer preferably above its glass transition temperature (Tg) followed by heat setting. This induces molecular orientation of polymer chains in radial and axial directions. The stress is applied to the polymer tube by expanding it radially by blow molding and by stretching the tube axially by applying axial load that result into biaxial orientation of polymer molecules. The tube is heated to desired temperature by heating the mold. Radial deformation of the tube is achieved by pressurizing the tube in the mold with inert gas under pressure. The degree of radial deformation is defined as ratio of outside diameter of tube after expansion and original internal diameter of tube. This ratio may vary between 1 and 20 or narrowly between 2 and 6. Degree of axial deformation is defined as ratio of lengths of tube after and before deformation. Temperature and degree of deformation affect crystallinity which in turn is dependent on crystallinity of the tube before deformation. The patent describes laser cutting of the deformed tube to get the scaffold structure of the stent.
U.S. Pat. No. 8,501,079 discloses method for fabricating a stent from PLLA tube; radially and axially expanding the tube inside a mold while the tube is heated to a processing temperature; wherein the processing temperature is 84° C. The radial and axial expansion percentages are 400% and 20% respectively to produce an expanded tube having an increased mechanical strength, fracture toughness and homogeneity in mechanical properties over the wall thickness of the expanded tube; and forming the stent from the expanded tube. The radial expansion of the tube is achieved at a pressure of 110-140 psi.
US 2013/0187313 discloses a method for fabricating stent comprising providing a PLLA tube disposed within a cylindrical mold; heating the mold and the tube to a tube deformation temperature (80° C. to 115° C.) with a heat source translating along the cylindrical axis of the mold and the tube, wherein the heat source translation rate is between 0.2-1.2 mm/sec; increasing the pressure inside the tube; allowing the increased pressure in the tube (110-140 psi) to radially expand the tube against the inner surface of the mold, wherein the radial expansion propagates along the cylindrical axis of the mold and tube as the heat source translates along the cylindrical axis, applying a tensile force to the tube along the cylindrical axis during the radial expansion to axially elongate the tube during the radial expansion, wherein the percent radial expansion is 300-500% and the percent axial elongation is 100-200%; and forming a stent pattern in the axially expanded and radially deformed tube.
EP1973502 reports a stent comprising a deformed spherical radio opaque marker disposed in a depot in a portion of the stent, the marker being coupled to the portion at least partially by an interference fit between an expanded portion of the marker and an internal surface of the portion of the stent within the depot, wherein the marker comprises sufficient radio opacity for ease of imaging by normal imaging techniques. Gaps between the deformed marker and the internal surface are filled with a polymeric coating material.
US 2011/0066222 describes method of forming a stent from PLLA tubular polymer that is deformed in a blow mold. Desired polymer morphology resulting in improved stent performance is obtained with axial expansion ratio from about 10%-200%, preferably 20% to 70%, a radial expansion ratio from about 100%-600%, preferably 400% to 500%, axial deformation propagation at or about 0.3 mm/min, selected expansion pressure of about 50 to 200 psi, preferably 130 psi and expansion temperature about 100° F. to 300° F. preferably less than 200° F. Heating is done by a moving heat source outside the mold. The heat source is moved at the rate of 0.1-0.7 mm per minute. The stent may be made of PLGA, PLLA-co-PDLA, PLLD/PDLA stereo complex, and PLLA based polyester block co-polymer containing a rigid segment of PLLA or PLGA and a soft segment of PLC or PTMC.
None of the prior art mentions the design and manufacturing method for making polymer stent with low strut thickness (less than 130 μm, preferably 100-110 μm thickness). The polymer stents have potential shortcomings compared to metal stents of the same dimensions viz. lower radial strength and lower rigidity of the polymer stents compared to metallic stents. Lower radial strength potentially contributes to relatively high recoil of polymer stents after implantation into an anatomical lumen. Another potential problem with polymer stents is that struts can crack or fracture during crimping, delivery and deployment, especially for brittle polymers. Due to these shortcomings, the strut thickness of the polymer stents is always kept higher compared to the metallic stents with same radial and fatigue strength.
In conclusion from “ISAR STEREO Trial” and “ISAR STEREO-2 Trial” as described above, it was established that the use of thinner strut device is associated with a significant reduction in angiographic and clinical restenosis after coronary stenting.
Thus, there is a continuing need for identifying right polymer and manufacturing and fabricating methods for polymeric stents of right scaffold design that impart sufficient radial strength, fracture toughness, low recoil and sufficient shape stability at low strut thickness. Additional advantage of a stent with low strut thickness is its lower profile after it is crimped on the balloon of the delivery catheter and more flexibility. Making a stent with desired low strut thickness starts with choosing a right polymeric material. The polymeric material then undergoes a number of process steps like drawing the tube from this material, modifying the mechanical properties of the tube, making the stent from this tube with right scaffold design, crimping the stent on the balloon of the delivery catheter and sterilization of the assembly.
The tube of the chosen polymer may be formed by extrusion or molding process under controlled conditions to achieve desired properties of the tube. Processing conditions that affect tube properties mainly include draw down ratio during extrusion, temperature at which the tube is extruded (relative to glass transition temperature and melting point of the polymer) and tube diameter.
Mechanical properties of a polymer can be modified by applying a stress. The stress alters the molecular structure and/or morphology of the polymer. The degree and rate of changes in mechanical properties depend on the temperature at which the stress is applied and degree of deformation the polymer (the tube in this case) undergoes due to application of the stress. The stress can be applied to the polymer tube in radial and axial directions to modify the crystalline morphology and polymer chain orientation in controlled manner to achieve a desired combination of strength and fracture toughness along axial and radial directions. Combined with right scaffold design, the strut thickness can be reduced while maintaining high fatigue and radial strength and keeping recoil under control. At the same time, it is necessary to achieve desired degradation rate of the polymer such that the stent retains adequate mechanical strength till the healing process of the lumen is completed and the stent eventually disappears from the implant site. The processing of the tube in this way changes the crystallinity of the polymer which in turn influences the degradation rate of the polymer. Amorphous polymer degrades faster than crystalline polymer but it is mechanically weaker than the crystalline polymer. Hence, a balance is required to be achieved in processing of the tube such that the stent has right combination of mechanical strength and degradation rate.
In the light of the above there is a need in the art to develop a biodegradable polymer stent with thin struts from a bioabsorbable polymer with adequate fatigue and radial strengths and a method to manufacture thereof. The process begins with choosing right grade of polymer and setting extrusion process to get desired properties of extruded tube. The grade of the polymer is characterized broadly by its molecular weight, glass transition temperature (Tg), crystallinity (Xc), molecular structure, and stereo isomerism. Further processing of the tube into stent and the scaffold design of the stent structure should be such as to achieve desired mechanical properties of the finished stent. The processing of the tube includes application of stress to the tube, laser cutting, cleaning of the cut stent, radio opaque marker deposition, heat treatment, drug coating, crimping and sterilization.
The mechanical properties are largely dependent on polymer characteristics like average molecular weight and molecular weight distribution. These characteristics undergo change at each processing stage. Hence it is necessary to check these characteristics at each process stage and devise a process that results in high mechanical strength of the finished stent.
Sterilization of the stent is done by e-beam radiation and this step requires special attention. The e-beam radiation causes degradation of polymer and thus has a significant effect on the average molecular weight of the bioabsorbable polymer and hence its mechanical properties. The current inventors have studied the effect of e-beam radiation on polymer over a wide range of e-beam doses and found that reduction in the e-beam dose improves the mechanical strength of the polymer. Normal dose of e-beam for effective sterilization is more than 20 kGy. The dose can be reduced to some extent by adding stabilizer/s in the polymer matrix. This stabilizer should be biocompatible and should not create any adverse clinical effect.
Therefore, one of the objectives of the instant invention is to get effective sterilization with e-beam dose considerably lower than 20 kGy without use of any additive.
Accordingly, the objective of the invention is to provide a biodegradable/bioabsorbable stent made from a bioabsorbable polymer with thin struts (thickness 130 μm and less, preferably with thickness of 100-110 μm) which has adequate fatigue strength, radial strengths and low recoil and a method for manufacture thereof, for which protection is sought.